1. Field of the Invention
The present invention relates generally to axial-flow blood pumps suitable for permanent implantation in humans as a chronic ventricular assist device.
2. Related Art
An effective and reliable axial-flow blood pump can provide mechanical circulatory support (MCS) to thousands of patients each year. An estimated 4.8 million Americans suffer from congestive heart failure (CHF), a clinical syndrome that involves ventricular dysfunction and ultimately a reduction in cardiac output. A reduction in cardiac output leads to poor perfusion, fluid accumulation, and activation of salt-water retention mechanisms. Statistics from the American Heart Association indicate that approximately 400,000 new CHF cases are diagnosed each year in the United States, and an estimated 40,000 cardiac failure patients die each year due to CHF.
If short-term medical intervention, whether surgical or through aggressive medications, is not successful, then many of these heart failure patients become candidates for cardiac transplantation. Due to a limited number of donor hearts available each year (˜2,500), CHF patients often require MCS as a bridge-to-transplantation and many such patients may fail to survive awaiting a donor organ. Ventricular assist devices (VADs) or mechanical blood pumps have proven successful in bridge-to-cardiac transplant support of patients suffering from end-stage heart failure and encourage the belief that long-term MCS or destination therapy is possible. An estimated 35,000 to 70,000 cardiac failure patients could benefit from long-term MCS each year.
Despite the tremendous need for an effective ventricular assist pump, prior MCS devices have not been entirely successful due to several limiting factors. Particular concerns and design limitations of current blood pump designs include: 1) component durability and lifetime; 2) blood clotting or thrombosis due to flow stasis within the pump in secondary flow areas, wash ports, and recirculation regions, and platelet activation in regions of high shear stress; 3) blood trauma or hemolysis that may occur when blood contacts mechanical bearings or foreign surfaces, and when blood is subjected to higher than normal shear conditions due to rotating components; 4) percutaneous leads which are needed for the motor and bearing control systems and other support lines; 5) pump geometry (size, shape, and weight) for ease of implantation, patient mobility, avoidance of graft tears; 6) high cost of the pump; and 7) high power demand that requires a large power supply.
Many early blood pump designs were constructed in a pulsatile configuration. These required the explantation (removal) of the diseased native heart and the replacement by the pulsatile mechanical blood pump. These pulsatile pumps, however, have proven to be very complicated mechanically, are relatively large and have relatively short mechanical lifetimes. An alternative design using a rotary pump leads to the use of the mechanical heart as a ventricular assist device which does not require explantation of the native heart. The rotary pumps have smaller size and better mechanical reliability. Such a pump aids a patient's heart by pumping additional blood in parallel with a diseased heart. The rotary blood pump may be connected to the patient's heart in a left-ventricular assist configuration, a right-ventricular assist configuration, or a biventricular assist configuration. For instance, if the left-ventricular configuration is adopted, the rotary pump is connected to receive flow from the left ventricle of the heart and return it to the aorta. Generally the rotary pump includes a stator (housing) having an inlet and outlet port, an impeller positioned within the stator and having impeller blades to create the pumped blood flow, a motor for rotating the pump and a suspension system. The blood enters the inlet of the stator and is pumped by the rotating impeller through the housing to the outlet, and back into the patient's circulatory system.
There are two primary configurations that are used for rotary blood pump configurations: axial flow and centrifugal flow. In the axial flow configuration, the pump configuration is similar to a cylinder with inlet flow port at one end and exit flow port at the other end. The centrifugal flow configuration is similar to a circular disk with an inlet flow port at the center of one side of the disk, oriented perpendicular to the plane of the disk, and a tangential exit flow port at the periphery of the disk, in the plane of the disk.
Studies have shown several problems with poor rotary blood flow path design in both axial flow and centrifugal flow blood pumps including those with magnetic suspension. One of these problems is stagnation resulting in thrombosis or clotting. If the flow undergoes a low or zero velocity region, it may experience thrombosis or clotting, where blood resides on the pump structure. Such low or zero velocity regions are usually found in secondary blood paths in the pump. As the thrombosis builds up, a section or large clot may break off and embolize in the blood stream. If the clot occludes a blood vessel that enters the brain or other sensitive area, very serious conditions may develop, such as profound organ dysfunction, such as seizure or severe brain damage. Another problem is hemolysis, where blood is exposed to high shear stresses in the rotary pump, usually near the impeller blades which move at relatively high speed, and which may cause direct or delayed damage to the circulating blood. As the impeller applies forces to the blood, regions of turbulence and/or jet formation, can occur in poorly designed devices.
Many rotary blood pump designs have been created to overcome these bearing problems with their use as ventricular assist devices. It is desired to have a bearing system with an expected operating lifetime of 10 to 20 years, if possible. Generally, these bearings fall into three categories: mechanical, hydrodynamic or magnetic bearings.
Some rotary blood pumps have mechanical or hydrodynamic bearings or hydrodynamic suspensions. For example, see U.S. Pat. Nos. 6,609,883 and 6,638,011. Other rotary blood pumps have a combination of hydrodynamic bearings and permanent magnet bearings. For example, see U.S. Pat. Nos. 5,695,471; 6,234,772; 6,638,083 and 6,688,861.
One type of rotary blood pump has mechanical bearings which require a lubricant flush or purge with an external lubricant reservoir for lubricating the bearings and carrying away heat. For example, see U.S. Pat. Nos. 4,944,722 and 4,846,152. There are many disadvantages to this type of pump. The percutaneous supply and delivery of the lubricant purge fluid degrades the patient's quality of life and provides a high potential for infection. Seals for the external lubricant are notoriously susceptible to wear and to fluid attack which may result in leakage and the patient having a subsequent seizure. Also, an additional pump is needed for delivery of the lubricant to the bearing, and if it fails the lubricated bearing freezes. Finally, the mechanical bearings have a finite wear life, usually of a few years, and need to be replaced due to the bearing wear.
There are axial flow rotary pumps with ceramic bearings presently under clinical trials. It is not known how long these bearings might last but expected lifetimes based upon other applications are in the range of 2 to 5 years. Also, there have been reported cases of thromboembolism in some patients. This has occurred while the patients are being anticoagulated.
Rotary pumps have been developed with magnetic suspension to overcome the earlier need for an external purge of lubricant or ceramic mechanical bearings. Utilizing a magnetically suspended impeller eliminates direct contact between the rotary and stationary surfaces, such as found in mechanical bearings. For example, see U.S. Pat. Nos. 5,326,344 and 4,688,998. Expected operating lifetimes of magnetic suspension systems range from 10 to 20 years. This type of rotary pump with magnetic suspension generally includes an impeller positioned within a housing, with the impeller supported and stabilized within the housing by a combination of permanent magnets positioned in the impeller and the stator, and with an electromagnet positioned within the stator. The impeller is rotated by a ferromagnetic motor consisting of a stator ring mounted within the housing, and electromagnetic coils wound around two diametrically opposed projections. The ferromagnetic impeller and the electromagnetic coils are symmetrically positioned with respect to the rotary axis of the pump impeller.
In magnetically suspended rotary blood pumps the gap between the stator and the impeller serves the competing purposes of allowing the blood to pass through, as well as assisting with the magnetic suspension and rotation of the impeller. For the blood flow, the radial gap is desired to be large for efficient blood pumping, but for efficient magnetic suspension, the radial gap is desired to be small. Because of the competing gap requirements, other prior art pumps often include a primary fluid flow region and a secondary magnetic gap. The primary fluid flow radial gap region is large enough to provide for hydrodynamically efficient flow without traumatic or turbulent fluid flow. The secondary magnetic radial gap allows for fluid therethrough which is small enough to provide for efficient magnetic levitation of the central hub, which can be either the stator or the impeller. Examples of pumps with a blood flow path including both a primary and secondary blood path can be found in U.S. Pat. Nos. 6,071,093; 6,015,272; 6,244,835 and 6,447,266.
Some prior art blood pumps include a permanent magnet thrust bearing which has a relatively large diameter thrust disk with permanent magnets having an alternating polarity configuration on both the stator and rotor components of the pump, but oriented in a radial configuration. It is believed that the large thrust disk obstructs the blood path and creates a tortuous blood path which is far from straight through the pump. The prior art pumps, however, include radially polarized permanent magnet configurations.
Various sensing techniques have been used to locate and control the rotor position of an active electromagnetic bearing. These techniques include eddy current or inductive, capacitive, and laser sensors, commonly used in industrial applications of electromagnetic bearings. Laser sensors cannot be used because they cannot “look through” the opaque blood. Eddy current and inductive sensors require a magnetic source in the stator and a magnetic target in the rotor as well as a magnetic path running between the stator and rotor used to sense the rotor position. Capacitance probes require an electrical path between the stator and rotor which is generally not feasible with blood pumps.
There are several problems associated with the use of an eddy current or inductive sensor types in the gap in between the stator and rotor of an implantable miniature blood pump. For example, these types of sensors rely on a clear magnetically un-obstructed pathway between the sensor “face” and the rotor surface. This means that the sensor body must be placed within the stator housing with its “face” perpendicular to the rotor surface. It is desired to avoid having any part of the body of the sensor placed within the fluid stream (blood) of the pump, yet placed it close to the rotor magnetic target. One problem with these types of sensors is the possible contamination of the blood stream if the soft iron or other non-biocompatible sensor faces are exposed to the blood. Generally this is solved with the use of some sort of thin biocompatible material covering the sensor face and target to avoid blood contact. Other problems are space constraints required for the sensor, and the energy budget required for such sensors.
In addition, such rotary blood pumps include a motor to rotate the impeller which, in turn, produces the needed blood flow. It is desirable that the motor have very long operating life, and operate fault free during a range of time up to one or more decades of service. Some pumps use brashness DC motors for this purpose which have a compact and efficient design without brushes. There is no brush wear so the expected life of such pumps is very long. One issue with such pumps is assuring that the motor will start in the necessary direction for the impeller to pump. The proper start up direction of rotation can be an important issue.
As noted above, magnetic suspensions often include at least one active control electromagnetic bearing axis. In a radial bearing configuration, the electromagnet can consist of a set of soft iron magnetic poles in a configuration in the stator arrayed circumferentially around the rotor with a clearance gap and imposing centering forces acting upon a soft iron placed in the rotor. The current in the coils must be controlled properly to achieve the desired centering purpose, allowing the rotor to properly operate in the clearance space. The current in the coils are adjusted by an automatic control system. In order to carry out the automatic force centering control method, there must be a sensor to sense the position or displacement of the rotor magnetic target, an electronic means of active control to adjust the control currents in the stator coils, such as electronic controller boards and power amplifiers to provide the power.
The control methods for determining the coil currents in the magnetic bearing of an active bearing can involve a manner of specifying how the coil currents are to be obtained. Active magnetic bearings can be composed of soft iron magnetic materials which have a significant limitation in that they are subject to magnetic saturation at a certain level, typically at about 1 Tesla. The active control method should take that fact into account. Also, the force exerted by an active magnetic bearing is a nonlinear function of both the magnetic gap and the current in the coils.
The most common method of dealing with both the magnetic material saturation and the nonlinearity is that of bias linearization, where a steady state bias current is imposed on each of the coil currents, which produces a magnetic flux in the soft magnetic iron poles of approximately one half of the magnetic saturation flux. Then a perturbation current is applied to produce changes in the coil currents associated with the poles on one side of the rotor relative to the other side. In turn, when these differences due to higher coil currents associated with the poles on one side produce higher magnetic forces on the rotor magnetic target compared to the other side, where lower coil currents associated with poles on the other side of the rotor produce lower magnetic forces acting on the rotor magnetic target. The rotor has a net force which is employed to center it in the clearance gap. The use of the bias linearization method, as just described, has a major disadvantage that it has relatively high power consumption in the coils, and generates large heating in the coils which may overheat the active magnetic bearing component of the rotary blood pump. Further, there are limitations on this type of bias linearization which prevent the full utilization of the magnetic force capacity of the active magnetic bearing.
Another issue with regard to proper centered operation of the rotary pump impeller is the unbalance in the rotor. Rotating devices are subject to mechanical unbalance as it is difficult to manufacture a perfectly balanced rotor. In addition, during operation within the patient over a long period of time, additional changes in unbalance may take place due to rotor component shifting, rotor rubbing, blood or blood products adhesion to the impeller surfaces, and other factors.
As the patient undergoes different levels of activity, the non-centering forces acting upon the rotor change. When the patient is active, such as in walking or climbing stairs, the magnetic current biasing levels in the active magnetic bearing are required to be high to provide high magnetic centering forces. However, when the patient is sitting quietly or sleeping, much less magnetic bearing centering force is needed. Higher bias current levels result in higher power consumption and higher heating.
In addition, centrifugal flow rotary blood pumps with magnetic suspension have been proposed. For example, see U.S. Pat. Nos. 6,074,180; 6,595,762, and 6,394,769. It has been found, however, that centrifugal flow pumps are not easily implantable in either animals or humans because the inflow and outflow cannulas are located at 90 degrees relative to each other and in separate planes. In addition, such centrifugal pumps require convoluted secondary blood flow paths as part of the design.